|Publication number||US7308074 B2|
|Application number||US 10/707,405|
|Publication date||Dec 11, 2007|
|Filing date||Dec 11, 2003|
|Priority date||Dec 11, 2003|
|Also published as||CN1626038A, DE102004059794A1, US20050129171|
|Publication number||10707405, 707405, US 7308074 B2, US 7308074B2, US-B2-7308074, US7308074 B2, US7308074B2|
|Inventors||Haochuan Jiang, David M. Hoffman|
|Original Assignee||General Electric Company|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (14), Referenced by (36), Classifications (12), Legal Events (4)|
|External Links: USPTO, USPTO Assignment, Espacenet|
The present invention relates generally to diagnostic imaging and, more particularly, to a CT detector having a reflector assembly with low cross-talk and high light output. In addition, the present invention relates to a reflector interstitially disposed between scintillators of a scintillator array that reduces cross-talk to improve CT image quality while simultaneously retaining high light output of the scintillators.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom.
Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
“Cross-talk” between detector cells of a CT detector is common. “Cross-talk” is generally defined as the communication of data between adjacent cells of a CT detector. Generally, cross-talk is sought to be reduced as cross-talk leads to artifact presence in the final reconstructed CT image and contributes to poor spatial resolution. Typically, four difference types of cross-talk may result within a single CT detector. X-ray cross-talk may occur due to x-ray scattering between scintillator cells. Optical cross-talk may occur through the transmission of light through the reflectors that surround the scintillators. Known CT detectors utilize a contiguous optical coupling layer(s), typically epoxy, to secure the scintillator array to the photodiode array. Cross-talk, however, can occur as light from one cell is passed to another through the contiguous layer. Electrical cross-talk can occur from unwanted communication between photodiodes.
Scintillator arrays typically incorporate a reflector layer or coating between adjacent scintillators to limit cross-talk between the scintillators. Generally, the reflector is formed of a material comprising chromium oxide or other types of optically absorbent material to absorb light transmitting across the separation boundaries between scintillators. Because chromium oxide operates as a good absorbent of light, the relative reflectivity of the reflector is reduced, which in some cases may be as much as 60%. As such, incorporating a reflector layer that includes chromium oxide, or similar material, a tradeoff in CT detector design is made between lower cross-talk and reflectivity. If the reflector layer is fabricated without chromium oxide or other optically absorbent materials, cross-talk between scintillators increases. Simply, implementing optically absorbent materials reduces cross-talk but lowers the reflectivity of the reflector.
Reduced reflectivity degrades low signal performance and increased cross-talk affects spatial resolution. Low signal performance is a function of noise generated in the CT detector. As reflectivity falls, the light output of the scintillator also falls. Noise, however, is relatively constant, therefore, decreases in light output increases the ratio of noise to functional light output. Additionally, the amount of cross-talk that may be attributed to scattered x-rays can be estimated to be about 50% of the total cross-talk in the CT detector. While the optically absorbent material is effective in reducing cross-talk associated with the transference of light between scintillators, the reflector typically has poor x-ray absorption characteristics and as such, does not eliminate the x-ray caused cross-talk that may occur between scintillators.
Therefore, it would be desirable to design a CT detector with reduced light and x-ray cross-talk characteristics to improve CT image quality without a sacrifice in light output for improved signal.
The present invention is directed to an apparatus for improving cross-talk reduction in a CT detector without significant reductions in scintillator light output. A method of manufacturing such an apparatus is also disclosed.
A multi-layer reflector for a CT detector is disclosed. The reflector includes an x-ray absorption component that is sandwiched between a pair of highly reflective components. Such a reflector is formed between adjacent scintillators of a CT detector so as to reduce cross-talk between adjacent scintillators as well as maintain a relatively high light output for signal detection. Moreover, the multi-layer reflectors may be disposed one-dimensionally or two-dimensionally across a scintillator array. A method of manufacturing such a reflector and incorporating same into a CT detector is also disclosed.
Therefore, in accordance with one aspect of the present invention, a CT detector includes a scintillator array having a plurality of scintillators and a reflector interstitially disposed between adjacent scintillators. The reflector includes a light absorption element disposed between a pair of reflective elements.
In accordance with another aspect of the present invention, a CT system is provided and includes a CT detector array having a scintillator array configured to illuminate upon reception of radiographic energy. The CT detector array further includes a reflector element disposed between adjacent scintillators of the scintillator array. Each reflector element includes a composite layer sandwiched between at least a pair of reflective layers.
According to another aspect of the present invention, a method of CT detector manufacturing is provided. The method includes the steps of providing a scintillator array of a plurality of scintillators and disposing a reflective layer between adjacent scintillators. The manufacturing method further includes the step of disposing a composite layer in the reflective layer.
Various other features, objects and advantages of the present invention will be made apparent from the following detailed description and the drawings.
The drawings illustrate one preferred embodiment presently contemplated for carrying out the invention.
In the drawings:
The operating environment of the present invention is described with respect to a four-slice computed tomography (CT) system. However, it will be appreciated by those skilled in the art that the present invention is equally applicable for use with single-slice or other multi-slice configurations. Moreover, the present invention will be described with respect to the detection and conversion of x-rays. However, one skilled in the art will further appreciate that the present invention is equally applicable for the detection and conversion of other high frequency electro-magnetic energy. The present invention will be described with respect to a “third generation” CT scanner, but is equally applicable with other CT systems.
Rotation of gantry 12 and the operation of x-ray source 14 are governed by a control mechanism 26 of CT system 10. Control mechanism 26 includes an x-ray controller 28 that provides power and timing signals to an x-ray source 14 and a gantry motor controller 30 that controls the rotational speed and position of gantry 12. A data acquisition system (DAS) 32 in control mechanism 26 samples analog data from detectors 20 and converts the data to digital signals for subsequent processing. An image reconstructor 34 receives sampled and digitized x-ray data from DAS 32 and performs high speed reconstruction. The reconstructed image is applied as an input to a computer 36 which stores the image in a mass storage device 38.
Computer 36 also receives commands and scanning parameters from an operator via console 40 that has a keyboard. An associated cathode ray tube display 42 allows the operator to observe the reconstructed image and other data from computer 36. The operator supplied commands and parameters are used by computer 36 to provide control signals and information to DAS 32, x-ray controller 28 and gantry motor controller 30. In addition, computer 36 operates a table motor controller 44 which controls a motorized table 46 to position patient 22 and gantry 12. Particularly, table 46 moves portions of patient 22 through a gantry opening 48.
As shown in
In one embodiment, shown in
Switch arrays 80 and 82,
Switch arrays 80 and 82 further include a decoder (not shown) that enables, disables, or combines photodiode outputs in accordance with a desired number of slices and slice resolutions for each slice. Decoder, in one embodiment, is a decoder chip or a FET controller as known in the art. Decoder includes a plurality of output and control lines coupled to switch arrays 80 and 82 and DAS 32. In one embodiment defined as a 16 slice mode, decoder enables switch arrays 80 and 82 so that all rows of the photodiode array 52 are activated, resulting in 16 simultaneous slices of data for processing by DAS 32. Of course, many other slice combinations are possible. For example, decoder may also select from other slice modes, including one, two, and four-slice modes.
As shown in
Referring now to
In one preferred embodiment, the thickness of the metal composite layer 86 is approximately 50-100 μm. In contrast, each reflective layer 88 preferably has a thickness of approximately 15-50 μm. The metal composite layers 86 are designed to absorb light that is transmitted from one scintillator to an adjacent scintillator thereby reducing, if not eliminating, optical cross-talk between the scintillators. Additionally, the metal composite layers are configured to absorb x-ray photons translating between scintillators. The amount as well as type of materials used in the metal composite layers defines the light as well as x-ray stopping power. However, one particular composite has been shown to absorb up to 50% of the x-ray photons between scintillators thereby reducing x-ray cross-talk by 50%. Given that optical cross-talk is typically 45% and x-ray cross-talk is typically 55% of the total cross-talk, with this exemplary composition and in accordance with the present invention, the total cross-talk of the scintillator array would be reduced by approximately 20% to 30% versus a conventional reflector. Additionally, the metal composite layer greatly reduces x-ray punch-through, e.g. by 60% or more.
Still referring to
Refining now to
In Stage B of the manufacturing technique, the substrate 94 undergoes one of a number of pixelating processes to define a number of scintillators 57 in the substrate 94. For example, the substrate 94 may be diced using a wire saw dicer or other dicing mechanism. Additionally, the individual scintillators 57 may be defined using ion beam milling, chemical etching, vapor deposition, or any of other well-known substrate cutting techniques. Preferably, the individual scintillators 57 are defined such that a gap 96 is formed between adjacent scintillators. Additionally, the scintillators 57 are preferably defined two-dimensionally across the scintillator substrate 94. Preferably, gaps 96 extend between individual scintillators 57 in both the x and z directions and have a width of approximately 100 to 200 μm depending on the requirement of geometric dose efficiency. The depth of the gaps depends on the stopping power desired and varies according to scintillator substrate composition.
Following formation or definition of the individual scintillators 57, a highly reflective material 89 is preferably cast onto the scintillators 57 and into the gaps 96 defined therebetween in Stage C. In one preferred embodiment, the cast filler 89 contains approximately 40% to 70% by weight titanium dioxide. However, one skilled in the art will appreciate that the cast filler 89 is not limited to an epoxy having titanium dioxide. Other highly reflective materials such as Ta2O5, HfO2, Bi2O3, and PbO, as well as other similar materials may also be used. While these materials typically do not have a reflectivity characteristic as high as titanium dioxide, these materials do have sufficient x-ray stopping power characteristics that assist in the reduction of x-ray cross-talk between scintillators. Moreover, one skilled in the art will appreciate that casting defines one particular means by which reflector material may be disposed between the scintillators. As such, the present invention contemplates other deposition processes including injection molding, for example.
Preferably, the highly reflective material 89 takes the form of a powder and is cast in gaps 96. As such, the powder is cured for a prescribed period. After curing, the top surface or portion of the scintillator array is machined to leave a top reflective layer 90 that has a desired thickness, e.g. 200 μm thick.
In Stage D, new gaps or channels 98 are created between scintillators 57 in the reflective material 89. Preferably, gaps 98 are created along both the x and z directions. Gaps 98 may be created using one of a number of cutting or dicing techniques as well as chemically-based etching processes. For example, gaps 98 may be formed using a wire saw or machining laser. Chemical etching, ion beam milling, as well as other semiconductor fabrication processes may also be implemented. In the example of a laser, a ND:YAG laser, CO2 laser, or an AR+ laser, or semiconductor laser may be used. In this example, the laser beam is focused on the center or middle of the reflective material disposed between the scintillators 57 and the width of the cut is adjusted so that a desired gap or channel width 98 results following the cutting process.
Wire saw dicing may also be used to machine gaps 98 in the reflective material 89 disposed between scintillators 57. For example, a wire having a diameter of 70 μm or less may be used to cut the desired gaps 98. In this regard, the wires are positioned on a spool (not shown) with a desired pitch. A mechanical fixture is then used to accurately position the wires and spool.
It is contemplated that at least two different types of wires may be used. That is, a metal wire with grinding media slurry feeding with the wires may be used. In this regard, the wires pass through the reflective material 89 and create the desired gaps. The grinding media may be diamond, SiC powder, alumina, and other well-known grinding media material. Preferably, the grinding media power has a grid size of 1,000 to 3,000 mesh. Another possible solution is to use a metal wire embedded with diamond or SiC media. OD (Outer Diameter) dicing saw may also be used. Regardless of the method, means, and mechanism to generate gaps 98, in a preferred embodiment, the thickness of the resulting reflective coating on the surface of each scintillator is approximately 15 to 50 μm.
Following formation of gaps 98 in the reflective material 89 between scintillators 57 so as to form a pair of separated reflective layers 88, a metal powder composite 86 is deposited into each gap 98 during Stage E. Preferably, the metal powder composite includes a high-Z metal such as tungsten or tantalum and is selected because of its high x-ray stopping power. Preferably, the metal powder of metal powder composite 86 has particular size of 0.5 to 5 μm. A low viscosity polymer such as epoxy, EpoTek® 301, polyurethane, or other low viscosity polymer is selected as a binder for metal powder composite 86. EPOTEK is a registered trademark of Epoxy Technology Inc. of Billerica, Mass. In this regard, 40% to 60% by volume of the metal powder is preferably homogeneously mixed with a liquid polymer. The mixture or composition 86 is then cast into gaps 98 created in the reflective material 89 of reflective layers 88. After casting, the mixture 86 is allowed to cure.
One skilled in the art will appreciate that other methods or techniques may be used to deposit the metal layer composition 86 between pairs of reflective layers 88. For instance, the high-Z metal particulars may be coated with an adhesive binder material such as a thermoplastic polymer coating. The coated metal particulars would then be cast into the gaps 98 with a small amount of solvent such as alcohol. The solvent may then be vaporized whereupon the resultant material is heated to melt the thermoplastic coating that will bind all the particulars together as well as serve as an adhesive between scintillators 57. Another method includes coating the high-Z particles with tungsten or with low temperature solder film. The solder film is then melted after being cast into the gap. After the film is formed, the scintillator array is ground or milled on the top surface to remove extra material of the metal composite and reflective material. Preferably, the top reflector 90 has a thickness of approximately 50 to 200 μm to maximize light output while minimizing x-ray attenuation.
Once the metal composite layer 86 interstitially disposed between pairs of reflective layers 88 is allowed to cure, the scintillator array is then machined at Stage F into a final and desired dimension. Additionally, the bottom surface 99 of the scintillator substrate is machined or ground to remove extra scintillator material and to attain a final and desired thickness. For example, depending on the type of scintillator being fabricated, the final thickness ranges from approximately 1.5 to 3 mm. The machined surface may then be optically coupled to a photodiode in accordance with well-known CT detector fabrication assembly.
Referring now to
The present invention has been described with respect to fabricating a multi-layer reflector disposed between scintillators of a CT detector for a CT-based imaging system. Further, fabrication of a rectangular shaped scintillator has been described. However, the present invention contemplates additional patterns or shaped cells being fabricated and a multi-layer reflector being disposed between scintillator cells. Additionally, the present invention has been described with respect to reflectors that are cast along one dimension, i.e., the z-axis. However, the reflectors may be formed using the aforementioned methods of manufacturing along an x and z axis thereby rendering a “checkerboard” full two-dimensional (2D) arrangement of reflectors. The present invention may also be implemented to create a partial 2D array of reflectors.
Therefore, in accordance with one embodiment of the present invention, a CT detector includes a scintillator array having a plurality of scintillators and a reflector interstitially disposed between adjacent scintillators. The reflector includes a light absorption element disposed between a pair of reflective elements.
In accordance with another embodiment of the present invention, a CT system is provided and includes a CT detector array having a scintillator array configured to illuminate upon reception of radiographic energy. The CT detector array further includes a reflector element disposed between adjacent scintillators of the scintillator array. Each reflector element includes a composite layer sandwiched between at least a pair of reflective layers.
According to another embodiment of the present invention, a method of CT detector manufacturing is provided. The method includes the steps of providing a scintillator array of a plurality of scintillators and disposing a reflective layer between adjacent scintillators. The manufacturing method further includes the step of disposing a composite layer in the reflective layer.
The present invention has been described in terms of the preferred embodiment, and it is recognized that equivalents, alternatives, and modifications, aside from those expressly stated, are possible and within the scope of the appending claims.
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|U.S. Classification||378/19, 250/370.09|
|International Classification||H01L31/09, G01N23/04, A61B6/03, G01T1/20, G01T1/24|
|Cooperative Classification||A61B6/4085, A61B6/032, G01T1/2002, A61B6/4411|
|Dec 18, 2003||AS||Assignment|
Owner name: GE MEDICAL SYSTEMS GLOBAL TECHNOLOGY CO. LLC, WISC
Free format text: ASSIGNMENT OF ASSIGNORS INTEREST;ASSIGNORS:JIANG, HAOCHUAN;HOFFMAN, DAVID;REEL/FRAME:014205/0987
Effective date: 20031208
|Apr 20, 2004||AS||Assignment|
Owner name: GENERAL ELECTRIC COMPANY, NEW YORK
Free format text: ASSIGNMENT OF ASSIGNORS INTEREST;ASSIGNOR:GE MEDICAL SYSTEMS GLOBAL TECHNOLOGY COMPANY, LLC;REEL/FRAME:016212/0534
Effective date: 20030331
|Jun 13, 2011||FPAY||Fee payment|
Year of fee payment: 4
|Jun 11, 2015||FPAY||Fee payment|
Year of fee payment: 8